Fabrication of Scaffolds for Bone-Tissue Regeneration - MDPI
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materials
Review
Fabrication of Scaffolds for Bone-Tissue Regeneration
Petra Chocholata , Vlastimil Kulda and Vaclav Babuska *
Department of Medical Chemistry and Biochemistry, Faculty of Medicine in Pilsen, Charles University,
Karlovarska 48, 301 66 Pilsen, Czech Republic; petra.chocholata@lfp.cuni.cz (P.C.);
vlastimil.kulda@lfp.cuni.cz (V.K.)
* Correspondence: vaclav.babuska@lfp.cuni.cz; Tel.: +420-377-593-281
Received: 30 November 2018; Accepted: 11 February 2019; Published: 14 February 2019
Abstract: The present article describes the state of the art in the rapidly developing field of bone
tissue engineering, where many disciplines, such as material science, mechanical engineering, clinical
medicine and genetics, are interconnected. The main objective is to restore and improve the function
of bone tissue by scaffolds, providing a suitable environment for tissue regeneration and repair.
Strategies and materials used in oral regenerative therapies correspond to techniques generally used
in bone tissue engineering. Researchers are focusing on developing and improving new materials
to imitate the native biological neighborhood as authentically as possible. The most promising is a
combination of cells and matrices (scaffolds) that can be fabricated from different kinds of materials.
This review summarizes currently available materials and manufacturing technologies of scaffolds
for bone-tissue regeneration.
Keywords: bone tissue engineering; scaffolds; regenerative medicine; stem cells; hydrogels
1. Introduction
Tissue engineering is a relatively new and a very multidisciplinary field. It interconnects many
disciplines, such as materials science, mechanical engineering, clinical medicine and genetics [1].
The main objective of tissue engineering is to restore and improve the function of the tissues by
preparing porous three-dimensional scaffolds, and seeding them with cells and growth factors [2].
These three things (scaffolds, cells, growth factors) are known as “the tissue-engineering triad”, and
this system is set up in an appropriate environment in a bioreactor [3,4]. The term “tissue engineering”,
where engineering and the life sciences are interconnected, was introduced in 1988 in the National
Science Foundation workshop as “the application of principles and methods of engineering and life
sciences towards the fundamental understanding of structure–function relationships in normal and
pathological mammalian tissues and the development of biological substitutes to restore, maintain or
improve tissue function.” Langer and Vacanti used this term in a review article published in Science in
1993 [5].
Tissue engineering uses different techniques and different systems. The simplest approach is
delivery of suitable signal molecules (tissue-inducing substances) to the right place. Most techniques
rely on cells or cell substitutes, which can replace non-functional cells. The challenges are to preserve
their function and immunological rejection. The most promising are combinations of both cells and
matrices, as a matrix can be used in different kinds of materials as a matrix, and there is a tendency to
investigate bioactive materials that mimic the native biological neighborhood [6]. This system could be
implanted or used as an extra-corporeal mechanism, where cells placed on or within matrix systems
can be open or closed (encapsulated). An open system is fully integrated into the body. The cells are
anchored in a matrix and the system is implanted into the body. Use of immunosuppressive drugs
or autologous cells is a way of avoiding immunological rejection. A closed system is protected from
Materials 2019, 12, 568; doi:10.3390/ma12040568 www.mdpi.com/journal/materials
Materials 2019, 12, 568 2 of 25
the patient’s immune system by a membrane that isolates the cells but allows nutrients and waste to
permeate [5].
Strategies and materials used in oral regenerative therapies correspond to techniques generally
used in bone tissue engineering.
2. Bone Tissue Engineering
Bone tissue engineering could be a way of repairing bone defects generated from various causes.
This field comprises three main parts in vivo or in vitro—seeding cells, growth factors and scaffold
materials [7].
The main target of tissue engineering is to simulate natural behavior. The big advantage of bone
tissue is its natural self-repairing, remodeling and regeneration. There is an aim to produce scaffolds
able to provide regenerative signals to cells. Scientists try to develop a way of producing scaffolds made
of biomaterials that mimic those found in the natural environment, with multi-functional properties,
such as improving cell adhesion, proliferation, and differentiation [8].
Bone transplantation is the second most common type of tissue transplantation following blood
transfusion. Due to the aging of the population, it is expected to be increasingly in demand [9].
Bone tissue engineering combines biomaterials and cells. In the field of bone regeneration, emphasis
has been placed on stem cells [10], particularly in connection with osteoblasts.
Based on principles of modern tissue engineering, craniofacial tissue engineering emphasizes
craniomaxillofacial applications and aims to develop biomaterials for regeneration of oral and dental
tissues, such as bone, dentin, cementum, periodontal ligaments, mucosa, and salivary glands [11].
Pulp cells were tested on a synthetic polymer scaffold in vivo and in vitro to regenerate the pulp, and it
was the fundamental research for pulp-dentin tissue engineering [12].
Regenerative endodontics (RE), a new part of regenerative medicine, aims to treat infection of the
dental pulp that can lead to inflammation and even to tissue necrosis. Root canal system treatment
may be a common way, but it could lead to re-infection and even to tooth loss at the end. RE tries to
discover alternative methods of pulp and dentin regeneration. Cell therapy is becoming a pivotal part
of RE, and human permanent dental pulp tissue stem cells (DPSCs) could be a promising source of
stem cells. DPSCs can be accessed easily, and they could be able to differentiate into various lineages
(e.g., fibroblast, nerve cells, endothelial cells and odontoblasts) to create new connective tissue [13].
An easy source of DPSCs could be human third molars, due to their high level of clonogenicity and
proliferation and ability to form calcified colonies [14]. The main target of RE is to develop suitable
scaffolds, including antibiotic mixtures, cells, and growth factors, for pulp-dentin complex regeneration.
Scaffolds could be produced from polymers, both synthetic (e.g., poly(lactic) acid) and natural (e.g.,
collagen) by various technologies [15].
Guided bone/tissue regeneration (GBR) is the most well-documented technique of periodontal
regenerative therapy [16]. GBR, also called “membrane protected bone regeneration” [17], uses barrier
membranes in the treatment of alveolar ridge defects and promotes bone growth into tissue defects
adjacent to dental implants. Mineralization of the newly formed bone matrix at the GBR region could
develop earlier than at the bone-implant interface, and it causes local mechanical stability due to
regional variation of interfacial bone properties [16].
3. Structure and Properties of Bones
3.1. Architecture of Bones
The bone tissue is well organized from macro- to nano-scale structures (Figure 1) [10]. The bone
extracellular matrix (ECM) consists of organic components (22 wt %), inorganic crystalline mineral
components (69 wt %) and water (9 wt %). Organic components consist of type I collagen, also type
III and type IV collagen, and fibrin [18]. In addition, there are over 200 types of noncollagenous
matrix proteins (glycoproteins, proteoglycans, sialoproteins, etc.) [9]. Inorganic crystalline mineralMaterials 2019, 12, 568 3 of 25
components are represented by hydroxyapatite and calcium phosphate. Bone tissue contains the
largest amount of calcium in mammals and it can be treated as a ceramic-organic bio-nanocomposite
complex [9]. Organic components ensure flexibility, whereas inorganic components ensure strength
and toughness [18]. The mechanical, biological and chemical properties and functions of bones depend
on the irregular but optimized structure, making bone material heterogeneous and anisotropic, as can
be seen at different levels (Figure 1) [18,19].
Figure 1. Different length scales in hierarchically organized bone. The macrostructure creates the overall
bone shape and consists of trabecular (cancellous, spongy) bone, 50–90 vol % porosity and compact
(cortical) bone, less than 10 vol % porosity [20]. The microstructure (of about 10–500 µm) consists of
the Haversian system, osteons and single trabeculae). The sub-microstructure (of 1–10 µm) consists
of lamellae. The nanostructure (a few hundred nanometers—1 µm) consists of fibrillary collagen and
embedded minerals. The sub-nanostructure (below a few hundred nanometers of minerals) consists of
collagen, non-collagenous organic proteins, and fundamental structural elements.
Two major types of bone structure can be distinguished: trabecular and compact bone. Trabecular
bone is formed by a porous trabecular network and bone marrow filling a large inner space. Compact
bone is made from inorganic crystalline mineral with a very low number of osteocytes, blood vessels,
etc. Both types of bones are reinforced by collagen fibers. Age, anatomical site and bone quality
influence the mechanical properties, the most important of which are strength and elasticity. Porosity
and architecture affect the properties of trabecular bone. Compact bone is more resistant to longitudinal
stress than to radial, and to compression than to tension [18].
3.2. Osteoblasts, Osteocytes, Osteoclasts and Bone-Lining Cells
During life, two inseparable processes—bone resorbing by osteoclasts and bone formation by
osteoblasts—happen alongside remodeling of the skeleton with optimal mechanical integrity. Without
integrity, there can be bone loss, especially in the form of osteoporosis [21]. Bone modeling takes place
during growing up, as well as in adulthood, in which it maintains bending resistance and function.
A long-term process of bone remodeling replaces damaged bone with new bone and maintains
functions. Modeling and remodeling maintain formed bone and participate in repair of bone fracture.
It has been established that about 25% of trabecular and 3% of cortical bone are removed and replaced
every year [22].
Marrow stroma is important for regulation of hematopoiesis. Endosteum, a source of mature
osteoblasts in adulthood, comprising bone-lining cells, is important for the regulation of bone formation.
Periosteum consists of two layers, a fibrous layer formed by collagenous tissue and a cambium
layer containing a large number of cells. Cells in the cambium layer are activated during bone
regeneration and fracture repair. Osteocytes (osteocyte perilacunar matrix) are present together with
vasculature in a lacunocanalicular system (not mineralized), generating bone surface and participatingMaterials 2019, 12, 568 4 of 25
in the production of ECM proteins that are important to phosphate metabolism and mineralization.
An overview of bone ECM components can be seen in Figure 2 [21].
Figure 2. Matrix compartments of bone.
4. Materials
4.1. Bone Repair Biomaterials
Bone plays an important role in the homeostasis of minerals. The most important ones, phosphate
and calcium ions, are stored there and, when required, they can be released into the blood. Another
important function of bones is to ensure locomotion, load-bearing capacity and protection of the
internal organs of the body. Bone tissue is dynamic and highly vascularized, a process that continues
throughout an individual’s lifetime. Most fractures do not require any surgical intervention because of
the high regenerative capacity of bone, but, unfortunately, large bone defects and non-union fractures
do require it [20].
Treatment of degenerating tissues can be maintained by using an autograft or an allograft.
Although both methods are revolutionary, there can be many problems that tissue engineering
(regenerative medicine) needs to solve [3]. Autologous bone grafting, considered to be a gold standard
in the treatment of bone defects, is a process in which a graft is taken from one anatomic site and
implanted in another. Coming from the same individual, such a bone graft is integrated faster and
more completely. On the other hand, there are several disadvantages, e.g., blood loss, longer surgical
time, infection and a limited quantity of graft material [22]. The most commonly used autografts
are cancellous, cortical, bone marrow and vascularized bone grafts [23,24]. During allogenic bone
grafting, a graft is taken from one individual and transplanted to another. This type of graft can
be adapted to an appropriate form. Allogenic bone grafting can be divided into cancellous, cortical
and demineralized [22]. Bone grafts should satisfy the following requirements: osteoconductivity,
osteoinductivity and osseointegration [25]. Osteoconductivity means that the bone could grow on
the graft surface or down into pores, channels, or pipes, resulting in formation of cancellous bone in
structure. In that case, mesenchymal stem cells can grow passively. Osteoinductivity is the ability to
make pluripotent cells into bone-forming cell lineage. It is induced by growth factors that support
mesenchymal stem cells in differentiating into chondroblasts and osteoblasts. During the osteogenesis,
new bone is formed [23]. Osseointegration is related to direct contact between bone and the implant.
Favorable incorporation of a graft is influenced by many factors, such as type of bone graft, site of
implantation, etc. (Table 1) [24].Materials 2019, 12, 568 5 of 25
Table 1. Bone graft activity by type.
Graft Osteogenesis Osteoconduction Osteoinduction Mechanical Properties Vascularity
A UTOGRAFT
Bone marrow ++ +/− + − −
Cancellous ++ ++ + + −
Cortical + + +/− ++ −
Vascularized ++ ++ + ++ ++
A LLOGRAFT
Cancellous − ++ + + −
Cortical − +/− +/− ++ −
Demineralized − ++ +++ − −
+, ++, +++ = extent of activity: − = no activity, +++ = maximal activity.
There is an effort to produce “bioactive” materials that can be integrated with biological molecules,
in contrast to the past when “bio-inert” materials were designed [5]. Materials that can replace
autologous or allogenic grafts consist of bioactive ceramics, bioactive glass, biological or synthetic
polymers, and composites. With these materials it is easier to avoid problems with transplantation and
implantation, such as infection or insufficient adaptation to environmental stresses. The ideal premise
is that the material should be replaced by newly regenerated biological tissue at the same time [10].
4.2. Scaffolds
The term scaffold is used for three-dimensional (3D) biomaterial that provides a suitable environment
for cells to regenerate tissues and organs. The aim is to produce scaffolds that are able to provide
regenerative signals to the cells and to simulate natural behavior. Scientists try to develop ways
of producing scaffolds comprising biomaterials that are very similar to the natural environment,
with multi-functional properties [8], and which are efficient in terms of cost and clinical use [7].
The most important aspect is the structure of the scaffolds. Interconnected pores and high
porosity allow cell attachment to facilitate 3D regeneration of tissue, cell growth, proliferation,
and differentiation, diffusion of waste and the degradation products of scaffolds. The pore size
must be large enough to allow migration of cells, but small enough to allow the binding of cells to the
scaffold [7]. The degradation of the scaffold must last as long as the regeneration of the tissue [26].
Thus, an ideal scaffold for bone tissue would be osteoconductive, biodegradable and with proper
mechanical properties. It is necessary to be able to deliver cells, to produce the scaffold in irregular
shapes and, of course, to be commercially viable [27]. Tissue engineering aims to produce artificial
constructs as 3D scaffolds containing appropriate cells implanted directly in vivo to stimulate and
to direct formation of the new tissues. The human body is a sensitive system and materials for 3D
scaffolds must be biocompatible, easily sterilizable, and have good mechanical properties [7].
Many types of scaffolds are made as hydrogels. A hydrogel is a 3D flexible network of natural
or synthetic polymer that is insoluble in water [28], e.g., polyethylene glycol or alginate-based
hydrogels [10]. It can retain a large amount of water or biological fluid due to hydrophilic groups
such as –NH2 , –COOH, –OH, –CONH2 , –CONH–, or –SO3 H. Hydrogels can mimic living tissues
due to responding to changes in environmental conditions such as pH, temperature, and electric
field. They are suitable for delivery in a largely non-invasive way and for in situ gelling at body
temperature [29]. It can be applied in bone regeneration, and its big advantage is the possibility of cell
encapsulation and chemical biofunctionalization [10].
Hydrogels are used as matrices for tissue engineering due to their porosity. Cells can grow and
proliferate there, drugs can be released from them, nutrients and waste products can be diffused
through them. Another benefit is easy modification by adhesion ligands. Self-cross-linking hydrogels,
e.g., water-soluble chitosan and oxidized hyaluronic acid, are used for cell carriers because they do not
need any chemical cross-linking agents [30]. On the other hand, there are also some disadvantages:
due to their weak mechanical properties it is difficult to handle and sterilize them [31].
Classification of hydrogels can be based on different aspects (see Table 2) [28,29].Materials 2019, 12, 568 6 of 25
Table 2. Classification of hydrogels based on different aspects.
Cross-Linking
Physically Physical State Source Preparation Degradation
Chemically
(Self-Assembled)
chemical cross-linking freeze thawing solid
natural copolymeric biodegradable
grafting—chemical, stereocomplex
radiation formation semi- solid
radical polymeration ionic interaction
synthetic homopolymeric
condensation h-bonding
non-biodegradable
enzymatic maturation liquid
polymeration (heat-induced hybrid interpenetrating
high energy radiation aggregation)
4.3. Types of Materials Used
From a historical point of view, the only important property of the first-generation biomaterials
was their biocompatibility, while biointeractivity was the aim of the second-generation biomaterials.
Whereas the first generation was passive, the second generation can rather promote tissue regeneration.
Third-generation biomaterials are bioresponsive, e.g., they can activate genes to influence all aspects of
proliferation and differentiation of cells [11,32].
Materials currently used for bone tissue scaffold fabrication are inorganic materials and natural
or synthetic polymers. To enhance the mechanical properties of polymers, to utilize their excellent
characteristics and to increase tissue interaction, composites of polymers and ceramics have been
developed. It can be said that inorganic materials (bioactive glasses) are characterized by similar elastic
modulus (about 40 GPa) as cortical bone (about 20 GPa). On the other hand, synthetic polymers feature
lower values of strength (about 10 GPa) [33], and natural polymers even lower (about 70 MPa) [16].
Another important aim is to develop scaffolds with the ability to deliver specific drugs, such as
growth factors or antibiotics, thus improving bone ingrowth, bone healing, and the treatment of tissue
defects [34].
Natural materials, e.g., polysaccharides (starch, alginate, chitin/chitosan, hyaluronic acid
derivates) or proteins (soy, collagen, fibrin gels, silk), help cell adhesion and function. However,
immunogenicity may appear because of pathogenic impurities, and mechanical properties and
biodegradability can be less easy to control [35].
Synthetic polymers, e.g., poly(lactic acid) (PLA), poly(glycolic acid) (PGA) and their copolymers,
are more common in cell transplantation and scaffolds for tissue engineering because of their superior
mechanical properties and degradation rate control [35].
Inorganic materials, e.g., metals, bioactive glasses, tricalcium phosphate (TCP), hydroxyapatite
(HAp) and their combinations, are another group used in bone-tissue engineering because of their
similarity to the bone mineral phase [10]. Other types of bone materials have been developed (e.g.,
HAp-TCP biphasic ceramics, wollastonite) [36,37]. The main advantages and disadvantages of different
types of materials are summarized in Table 3.
4.3.1. Metals
Metals might be considered the oldest material used for implants [38]. The first recorded use
of metal implants was in Egyptian times [39]. The first metals used were aluminum, lead, gold and
silver [40]. Nowadays, titanium and its alloys are the most frequently used metallic biomaterials
for dental and orthopedic implants, as a result of their biocompatibility, non-toxicity and corrosion
resistance. Commercially pure titanium has excellent biocompatibility but relatively poor strength.
On the other hand, titanium alloys have superior strength, but contain ingredients that can be toxic or
allergenic [41]. Metal alloys are applied as joint replacements and fracture-fixation implants, because
of their good biocompatibility, corrosion resistance and strength [42,43]. Unfortunately, metals are not
biodegradable, so there is usually a requirement to remove metallic implants, especially in the case of
children [42].Materials 2019, 12, 568 7 of 25
Table 3. Types of materials used in tissue engineering and their advantages and disadvantages.
Type of material Advantages Disadvantages
Biocompatibility, non-toxicity and
Metals Not biodegradable
corrosion resistance
B IOCERAMICS
Bioactive glasses Improve differentiation and osteogenesis Low strength and brittleness
Bioactivity, biocompatibility, osteoconductivity,
Hydroxyapatite Brittle, very slow degradation
non-toxicity and non-inflammatory
Tricalcium phosphate Supports in vivo osteogenic differentiation Slow degradation, incompressible nature
N ATURAL POLYMERS
Collagen Enzymatic biodegradability Complexity of structure
Gelatin Biocompatible, biodegradable Poor mechanical properties
Support cell attachment, differentiation, and
migration, non-toxicity, non-allergenicity,
Chitosan Poor mechanical strength
mucoadhesivity, biocompatibility,
biodegradability and osteoconductivity
Biocompatibility, biodegradability, viscoelasticity,
Hyaluronic acid Very rapid degradation and water solubility
enzymatic biodegradability
Biocompatibility, easy gelling, easy chemical Non-degradable in mammals, sterilization
Alginate
modification causes degradation
Wide range of gelling and melting temperatures,
Agarose no need cross-linking agents, little inflammatory Poor cell attachment
response in vivo
S YNTHETIC POLYMERS
Degradation by bulk erosion, relatively
Degradation products can be excluded from
Poly(α-hydroxy acids) poor mechanical properties, hydrophobicity
the body
of the polymer surface
Poly(ε-caprolactone) Biodegradable, non-toxic, a low melting point Hydrophobicity, slow degradation
Excellent mechanical properties, Toxicity of degradation products (from
Polyurethanes
good biocompatibility aromatic diisocyanate component)
4.3.2. Bioceramics
Bioceramics include mechanically strong biomaterials, e.g., ceramic composites, amorphous
glasses and crystalline ceramics [42]. Bioactive glasses (BGs) and glass ceramics are biomaterials used
not only in bone-tissue engineering but also in orthopedics and dentistry [44]. The most commonly
used bioceramics in bone-tissue engineering are HAp, TCP and their composites [42]. Ceramic scaffolds
of HAp and TCP applied for bone regeneration are characterized by high mechanical stiffness, very
low elasticity, and brittleness. Due to their chemical similarity to native bone they show excellent
biocompatibility and facilitate good differentiation and proliferation of osteoblasts. On the other hand,
their brittleness can cause problems with mechanical loading and degradation-rate control [1].
Bioactive glasses (calcium and phosphate containing silica glasses) produce bioactive
hydroxyapatite after immersion in biological fluid and are able to bond to biological tissue. They can
improve differentiation and osteogenesis because they deliver silica ions that are necessary for
activating gene-transduction pathways [45]. It takes years to resorb bioactive glasses and bioceramics
with crystalline HAp. For better reabsorption, other calcium phosphates can be used, but their
disadvantages include low strength and brittleness, which are the characteristics that make it
impossible to use inorganic materials for load-bearing applications [10]. The most usual applications
of bioactive glasses are bone-filling materials, small bone implants, coating orthopedic implants and
dental applications [44]. The two main manufacturing processes used to make bioactive glasses are
melt-quenching and the sol-gel route. Before quenching oxides in water, they are melted at a high
temperature [42]. Bioactive glasses and ceramics are able to form a layer of active hydroxy carbonate
apatite (HCA), which connects to tissue upon implantation owing to their chemical and structural
similarity to the bone mineral phase. Ceramics can be modified to create an apatite layer in vivo in
protein-free simulated body fluids (SBF) [34].
Hydroxyapatite (HAp) (Ca10 (PO4 )6 (OH)2 ), as a major natural inorganic component of bone,
shows excellent bioactivity, biocompatibility, osteoconductivity, non-toxicity and non-inflammatory
characteristics. Synthetic HAp is white, whereas natural HAp can have various colors (brown, yellow,
green) [9]. Its mechanical properties are essentially influenced by the size of the HAp particles [46],
porosity, density, etc. [9]. HAp is very hard but brittle, with a very slow degradation rate in vivo,Materials 2019, 12, 568 8 of 25
and that is why it should be joined with natural or synthetic polymers to create scaffolds. On the
other hand, HAp is very beneficial for constructing bones, because it stimulates growth factors (e.g.,
bone morphogenic protein) and encourages alkaline phosphatase (ALP) in mesenchymal stem cells
(MSCs) [42].
Tricalcium phosphate (TCP) supports in vivo osteogenic differentiation of MSCs and is usually
used for the production of scaffolds. Injectable 3D scaffolds of beta-TCP (β-TCP), in combination with
alginate gel and also with type I collagen, are also available. Collagen scaffolds in combination with
TCP implanted in a rabbit femur bone showed better bone formation than collagen-HAp scaffolds [42].
4.3.3. Polymers
Two types of biodegradable polymers are used. The first are natural polymers and the second
are synthetic. Natural polymers such as polysaccharides (starch, alginate, chitin/chitosan, hyaluronic
acid derivatives) or proteins (soy, collagen, fibrin gels, silk) provide admirable cell attachment and
growth. On the other hand, they have many disadvantages, e.g., immune-response problems, and poor
mechanical properties [10]. The challenge in this field of material engineering is to produce natural
polymer-based scaffolds of sufficient quality and homogeneity [3].
Collagen is a basic component of animal tissues, such as bone, cartilage, tendons, skin, and blood
vessels [3]. Its polypeptide chain is very flexible and contains mostly glycine, proline, hydroxyproline
and lysine. The degree of flexibility is determined by the amount of glycine. Although about
twenty-nine types of collagen are known, the most widespread is type I [42]. As collagen is the
main part of ECM, it fulfils all the requirements for biomedical applications. Its principal advantage is
enzymatic biodegradability [47]. The process of biodegradability, as well as mechanical properties,
can be modified by cross-linking or by combination with inorganic compounds, e.g., HAp [42].
Collagen-HAp scaffolds have been produced with optimal pore structure for bone regeneration [3],
where 99% interconnectivity and excellent cell infiltration have been achieved. Scaffold derived from
mouse-bone marrow from mesenchymal stem cells was implanted into a mouse calvarial defect.
After three weeks, the defect was annealed and after several weeks, degradation of collagen-HAp the
scaffold was proven successful. Furthermore, it was confirmed that the osteogenic differentiation of
cells was improved by the addition of magnesium nanocrystals. Using human-derived bone MSCs
with this type of scaffold could be a way of treating osteochondral defects [42].
Gelatin is derived from collagen [48] by breaking the structure of the triple-helix into a single
strand [49]. This material is biocompatible as well as biodegradable because of its amino acids (such
as arginine-glycine-aspartic acid), which also encourage cell adhesion, migration, differentiation
and proliferation [48]. Although gelatin comes from collagen, its antigenicity is lower, but there are
still some information signals [50]. Change of its solution temperature causes gel formation, which
is utilized not only for wound dressing [51], but also for delivering a variety of drugs. The poor
mechanical properties of gelatin gels have led to the production of gelatin-based composite scaffolds
(e.g., ceramic-gelatin) for cartilage and bone repair [47].
Chitosan is one of the most widely used materials for scaffolds. In bone-tissue engineering,
it can be used alone or with other polymers or ceramics [52]. Chitosan has the same structure
as a non-collagen organic component of ECM glycosaminoglycan (GAG) [48]. This is a linear
polysaccharide consisting of D-glucosamine and N-acetyl-D-glucosamine linked by β (1-4) glycosidic
bond [47,53]. Chitosan can be obtained by deacetylation of chitin, which is an exoskeleton component
of crustaceans [54]. Free amine groups facilitate inherence through a positive charge for binding
with different negatively charged molecules (e.g., lipids, cholesterol, metal ions, proteins, etc.) [49].
The degree of deacetylation affects the crystallinity of chitosan, and 100% deacetylated chitosan is
highly crystalline [55]. Chitosan might support cell attachment, differentiation, and migration [48],
and because of its non-toxicity, non-allergenicity, mucoadhesivity, biocompatibility, biodegradability
and osteoconductivity, it can be applied as a dental, bone or cartilage implant, or as artificial skin.
Cross-linked chitosan can be used as a bandage. Chitosan nanofibers used for burn healing can protectMaterials 2019, 12, 568 9 of 25
against infection, absorb exudate, provide air access to a wound and help in the regeneration of skin
tissue [47]. It has been stated that chitosan promotes the growth of osteoblasts and the mineralization of
matrices [56], and for that reason it is extensively used as a sponge in bone-tissue engineering. Chitosan
sponge is a flexible, soft material with interconnected pores but with poor mechanical strength [47]
and therefore it is very often mixed with other natural polymers or bioceramics to obtain scaffolds.
3D scaffolds of HAp/chitosan-gelatin present a structure that is similar to human bone. The presence
of HAp improves mechanical properties, and the combination of natural polymers, ceramic material
and cells exhibits the effect of biomineralization after three weeks [57]. Chitosan-based composite
biomaterials are not only a very good option for cartilages and intervertebral discs, but also for gene
therapy in orthopedics.
Hyaluronic acid (HA) is abundant throughout the extracellular matrix of the human body. It is
composed of a linear glucosaminoglycan, where units of N-acetyl-D-glucosamine and D-glucuronic
acid are repeated [58]. Not only its biocompatibility and biodegradability but also its viscoelasticity
are convenient properties for using HA in biomedicine [47]. A very significant advantage of HA is its
enzymatic degradability by hyaluronidase [59], an enzyme produced by mammalian cells [47]. In view
of the very rapid degradation and water solubility of HA, it is advisable to cross-link it. HA hydrogels
can be produced through its chemical modification by mono- and polyvalent hydrazines [60], and
HA in this form is used for scaffolds. Modified HA can be used in the form of sponges, obtained
by lyophilization, for the culturing of human fibroblasts, chondrocytes and bone marrow MSCs [61].
Excellent adhesion and proliferation of human adipose-derived MScs on heparin-HA hydrogel) was
found. The heparin significantly increased cell adhesion compared to HA alone. When ADSCs were
encapsulated into heparin-HA hydrogel, a distinctive expression of hyaluronidase was observed.
Heparin-HA hydrogels with encapsulated stem cells without any modification could be a promising
system for application in biomedicine [59].
Alginate is a polysaccharide composed of β-D-mannuronic acid and α-L-guluronic acid connected
by (1-4) bond [47]. Its biocompatibility and ability to gel easily facilitate alginate which can be used in
many biomedicine applications [62]. Anionic polymer alginate can be obtained from seaweeds and
is already in commercial production [47,62]. Commercially used alginate presents higher purity and
therefore there is hardly any inflammatory response [62]. Alginate can be easily modified chemically,
and use is often made of this in its applications, e.g., wound-dressing materials, drug-delivery systems,
and implantation of protected living cells [63]. Tissue engineering uses alginate as a delivery system
via encapsulation techniques. The first encapsulation was used for rats’ Langerhans islets to transplant
cells to affect type I diabetes. Application in mammalian reproductive technologies was started
by encapsulation of bovine sperm [47]. In mammals, alginate is non-degradable because of lack
of enzymes, but degradability can be increased by its ionic cross-linking [62], and the degradation
can proceed not only in an acid (pH < 5.0) or basic (pH > 10.0) environment, but also in a neutral
environment in the presence of reducing compounds. Unfortunately, sterilization, e.g., by heating,
autoclaving, ethylene oxide or γ-radiation, causes degradation of alginate [63]. Alginate gels are
widely used in bone-tissue engineering for bone and even cartilage regeneration because gels can be
easily introduced into the body in a minimally invasive way and irregular shapes can be filled in.
Agarose is another polysaccharide that is useful for tissue engineering [64]. It consists
of repeating units of agarobiose (1,3-linked disaccharide of β-D-galactose and 3,6-anhydro-α-L-
galactopyranose) [65]. Agarose is often combined with other polysaccharides to form hydrogel
scaffolds [64] because of its poor cell attachment [65]. Agarose is a widely used compound not only in
tissue engineering and drug controlled release, but also as a gel for electrophoresis, chromatography
and, due to agar similarity, as a culture medium [66]. It can be obtained by extraction of agarophyte
seaweed (Rhodophyceae algae) cell walls [64]. Agarose is soluble in hot water and can be prepared
in the form of a thermal-reversible gel. Its temperatures for gelling and melting are from 30–40 ◦ C
to 80–90 ◦ C, depending on molecular weight, concentration and the number of its side groups [65].
The mechanism of agarose gelation depends on hydrogen bonds; helical structure is formed at firstMaterials 2019, 12, 568 10 of 25
and then the gel appears [67]. There is no need for cross-linking agents, because of the formation
of hydrogen bonds [65]. Due to the fact that mechanical properties can be adjusted, agarose can be
available for different applications in tissue engineering, especially in bone and cartilage scaffolds.
Agarose-based materials can create 3D hydrogel scaffolds that provide an appropriate environment
for cell growth including good permeability for oxygen and nutrients. By adding HA into agarose
composite, agarose/HA hydrogels were created and pore size, swelling ratio and thermal stability
increased. Little inflammatory response in vivo was reported. The HA content determines the rate
of degradation: the higher the HA content, the higher the rate of degradation [68]. Composites
containing HAp showed higher ALP activity, and 3% of micro-HAp was confirmed as the optimal
amount for calcification [69]. New Zealand rabbits were used for an animal experiment. HAp/agarose
gel composite was injected in the drilled holes, and the observation periods were 1, 2, 4 and 8 weeks.
Pure agarose gel was used for the control group. After 8 weeks, excellent bone formation was observed
by using micro-computed tomography analysis. After 4 weeks, little bone regeneration was observed
in bone defects containing pure agarose gel [70]. There is a need for more systematic research involving
agarose-based biomaterials, but these biomaterials will definitely find clinical applications [65].
Synthetic polymers are based on polyesters, such as polylactic acid (PLA), polyglycolic acid (PGA),
poly ε-caprolactone (PCL) or poly (lactic-co-glycolide) (PLGA) copolymers [47]. They can be produced
with a tailored structure, and degradation rate can be controlled quite easily [3]. Their disadvantage is
that there can be problems with their reduced bioactivity.
Poly(α-hydroxy acids) including poly L-lactic acid (PLLA), polyglycolic acid (PGA) and poly
D,L-lactic-co-glycolic acid (PLGA) copolymers, are the most widely used synthetic polymers for
3D scaffolds in bone-tissue engineering, but were historically used as resorbable surgical sutures.
According to need, the degradation rate can be adjusted from weeks to several years. The degradation
products can be excluded from the body as carbon dioxide and water [3]. To support cell distribution
and diffusion of nutrients, meshes, fibers, sponges and foams are produced as scaffold types.
There are several disadvantages to these materials. The first one is that these polymers degrade
by bulk erosion, which causes premature degradation [47]. Increasing amounts of acidic degradation
decrease pH, which can accelerate the degradation rate and eventually cause inflammation. Another
disadvantage of porous scaffolds is their relatively poor mechanical properties, especially in vivo [64].
It is possible to obviate these disadvantageous properties by preparing copolymers, composites of
polymer/bioactive ceramic [47], especially HAp, for bone-tissue engineering [64], or by adjustment
of molecular weight [47]. It has been reported that compact PLLA/HAp composites have good
osteoconductive properties, a better cell environment for seeding and growing [71], and surface
characteristics that are important for osteoblastic cells [47]. Degradation products do not change pH,
which was detected during 24 weeks of monitoring, and acidic products might act as buffers. Not least,
the mechanical properties were also improved [71].
Seeding and delivering the cells into porous scaffolds of PLA and PLGA is negatively affected by
the hydrophobicity of the polymer surface, which is another reason for combination with HAp and
bioactive glasses [47]. Porous polymers with a pore size of 100–500 µm, combined with bioceramic
particles, were found to be optimal scaffolds for bone-tissue engineering. The degradation rate,
as well as potential inflammation, can be affected by the composition of the polymers. Because of
possible toxicity of the residual solvent, it was found that the optimal production techniques were the
solvent-free method, gas foaming and rapid-prototyping [72].
Poly(ε-caprolactone) (PCL), a semi-crystalline, biodegradable [73], non-toxic in nature [74],
aliphatic polyester, with a low melting point (60 ◦ C) [47], can be used for easy production of scaffolds
for tissue engineering [75], bone and cartilage repair, surgical sutures, and drug-delivery systems [47].
Although it is more stable, cheaper and readily available, and in higher quality, than polyhydroxy
acids [76], the main disadvantages of PCL are its hydrophobicity, which is disfavorable for cell
attachment and infiltration, and its slow degradation, which can last up to 3 or 4 years [75].
Modification of its properties can be achieved by co-polymerization or blending with other polymers.Materials 2019, 12, 568 11 of 25
Co-polymerization directly changes chemical properties that influence other properties indirectly.
In comparison with blending, it changes physical as well as chemical properties and the biodegradation
rate [77]. Blending PCL with bioactive glass and bioceramics, such as HAp, can improve hydrophilicity
and bioactivity of scaffolds. It was found that a blend of PCL and magnesium phosphate (MP)
contained interconnected pores and achieved porosity of 73%. MP particles increased hydrophilicity
and the MP can be used for controlling the degradation rate [75]. Co-polymerization of ε-caprolactone
with methoxy poly(ethylene glycol) block copolymers for drug-delivery systems can be obtained and
hydrophilicity and even lipophilicity can be altered [73]. PCL can be combined with both natural and
synthetic polymers [76]. Scaffolds of PCL prepared by solid free-form fabrication (SFF) can differ in
internal architecture and porosity. Even the selective laser-sintering method (SLS) seems to be sufficient
for PCL scaffolds used in bone and cartilage tissue engineering. The mechanical properties of such
PCL scaffolds are within the lower range of trabecular bone. Scaffolds can sustain appropriate stress,
and SLS is an easy method for scaffold production [76].
Polyurethanes (PU), a major class of synthetic elastomers [78], are products of reaction of
molecules with two or more hydroxyl groups and molecules containing two or more isocyanate
groups. Two thermodynamic incompatible phases can be obtained [77]. Polyester chains comprise
the soft segments, while the hard segments are composed of polyurethane blocks on aromatic
isocyanates, providing non-biocompatibility due to the toxic degradation products. If polyurethanes
are designed to have chemical linkages, they can be degraded in the biological environment.
The toxicity of degradation products, the main disadvantage of polyurethanes, could be reduced
by using lysine diisocyanate (LDI) or other aliphatic diisocyanate (e.g., hexamethyl diisocyanate,
1,4-butandiisocyanate). Another way to prepare biodegradable and non-toxic polyurethanes can
be using star-shaped polyester prepolymer (from myoinositol). This type of the polyurethane was
implanted in guinea pig, where the biodegradability was demonstrated [78]. Polyurethanes are used
for fabrication of medical implants, especially for long-term implants [70] and biomedical products
such as cardiovascular catheters, diaphragms of blood pumps, coating materials for implantable
pacemakers, etc. [78].
4.3.4. Composite Materials
As we have seen above, there are many problems with scaffolds produced from a single
biomaterial (advantages and disadvantages listed in Table 3). The best solutions at present are
composite scaffolds, i.e., a combination of ceramics and polymer, or of synthetic polymers with
natural polymers [1]. Composite materials include a polymer phase with toughness and compressive
strength and an inorganic phase with bioactivity, which improves the mechanical properties and
degradation rate. That is why these materials are very similar to the natural structure of real bone.
Sol-gel processing is a technique that can create a polymeric network with inorganic components.
Unfortunately, their mechanical properties are not yet as good as the mechanical properties of bone [10].
Each phase has different properties, and tissue-engineering templates must interconnect the best
properties of both kinds [79].
Composite metal scaffolds can achieve an increase in biodegradability [80]. Magnesium (Mg) is
one of the most promising metals in combination with nutrient elements, strontium (Sr), calcium (Ca),
and alloys of Mg–Sr, Mg–Ca–Sr were produced and evaluated in vitro with bone marrow-derived
MSCs. The study showed that ternary Mg–Ca–Sr alloys had a higher degradation rate and better
cell adhesion compared to binary Mg–Sr alloys. Generally, Mg–1Sr, Mg–1Ca–0.5Sr and Mg–1Ca–1Sr
were recommended as the most suitable for in vivo studies in animals and also for clinical testing [80].
Another success of proliferation and osteogenic differentiation was achieved by preparing composites
of Sr, HAp and chitosan with human bone marrow MSCs. Improvement of attachment and proliferation
was also achieved for composites of nickel-titanium treated with sodium hydroxide, titanium-based
scaffolds, porous 3D injected iron-magnesium scaffolds, composite titanium-silica scaffolds with
complex geometry and stainless steel, titanium and cobalt chromium alloys [42]. Porous Nb–Ti–TaMaterials 2019, 12, 568 12 of 25
alloys can induce apatite in vitro and that is why they might be a new candidate for application
in bone-tissue engineering [43]. Many bio-inorganic ions (silicon, zinc, copper, lithium and cobalt)
incorporated into graft material for bone-tissue engineering proved equally beneficial for healing bone
defects, but further research is needed [22].
5. Manufacturing Technology of Composite Scaffolds
5.1. Methods of Scaffold Fabrication
Incorporation of advantageous properties allows the production of composite scaffolds [34].
There are several processing technologies that have been shown to produce porous 3D polymeric scaffolds
for bone tissue, especially used is solvent casting and particulate leaching, gas foaming, emulsion
freeze-drying, electrospinning, rapid prototyping and thermally induced phase separation [81].
5.1.1. Solvent Casting and Particulate Leaching
The solvent casting and particulate leaching technique is the most common and easy method,
where pore size and porosity (depending on the salt/polymer ratio) can be controlled. This technique
consists of dissolving polymer in an organic solvent, mixing it with water-soluble porogen, such as
salt (e.g., sodium chloride, sodium citrate) [82], and casting the resulting mixture into a mold [34].
The solvent evaporates or lyophilizates, and the polymer/porogen composite is leached into water.
The porosity of the scaffold is dependent on the amount of porogen, and the size of the pores depends
on crystal size. Waxy hydrocarbons [82] and gelatin particles [83] are alternative porogens. It was found
that a high interconnectivity of pores could be achieved at 70 wt % of porogen. The solvent-casting
method does not require any special equipment (flat sheets and tubes) [82]. Remaining toxic solvent can
cause denaturation of incorporated molecules, a decrease in the activity of bioinductive molecules [34],
and impossibility of adding pharmacological agents [82]. The polymer-ceramic 3D scaffolds obtained
feature-controlled pore interconnectivity and porosity at low porogen levels [34].
5.1.2. Gas-Foaming Process
The gas-foaming process is a technique where it is possible to avoid organic solvents, and pores
are created by gas expansion [84]. Carbon dioxide is used as a porogen gas [82], and this common gas
is low-toxic and non-flammable [84]. This method affords highly porous foam (pore size of 100 µm)
with a porosity of up to 93%, but the pore interconnection, especially on the surface, is low, at just
10%–30% [81]. Open porous scaffolds can be produced by using the gas foaming/salt leaching method.
A combination of ammonium bicarbonate salt and acid substances at increased temperatures causes
gaseous ammonia and carbon dioxide to be released. Macroporous structures with interconnected
pores of 100–500 µm can be obtained where cells can be seeded and their viability is high. This is a
simple and financially viable method of scaffold production [85]. This technique cannot be used for
hydrophilic and glassy polymers, owing to low solubility in CO2 , e.g., chitosan, but use of a co-solvent
such as ethanol or diluted acid could solve the problem [84].
5.1.3. Thermally Induced Phase Separation
Thermally induced phase separation (TIPS) allows production of porous anisotropic polymer
scaffolds that can be easily controlled and which have a low probability of defects for tissues such
as nerves, muscles, tendons, ligaments, the intestines, bones and teeth [86]. Their properties, e.g.,
pore morphology, mechanics, bioactivity and degradation rate, depend on polymer concentration and
the volume of the secondary phase fraction [81]. Generally, a polymer is dissolved in a solvent at a
high temperature, and a porous polymer scaffold is obtained by cooling the homogenous solution,
which causes phase separation, and a microporous structure arises after removal of the solvent [86].
Because of pore size (10–100 µm), however, this is not very suitable for seeding of the osteoblasts or
for bone-tissue growth. Therefore, a coarsening process is needed in the thermally induced phase inMaterials 2019, 12, 568 13 of 25
order to generate a pore size greater than 100 µm. The properties of scaffolds depend on polymer
concentration, solvent type and temperature gradient. The use of organic solvents is a disadvantage of
this technique, because solvent sublimation takes a long time [81].
Dimethyl sulfone (DMSO2 ) was recommended as a universal crystallizable solvent for polar
polymers. As a result of the use of DMSO2 , TIPS can be called a green method, because the solvent
can be recovered by recrystallization and sublimation [86]. Another type of solvent is PolarClean®
(sonicated Methyl-5-(diethylamino)-2-methyl-5-oxopentanoate), which is described as environmentally
friendly because it is non-toxic, water-soluble and biodegradable [87].
5.1.4. Solid Free-Form Fabrication Technique
The solid free-form fabrication technique (SFFT), also known as rapid prototyping (RP) or additive
manufacturing, belongs to the computer-controlled methods, such as shape- or size-solving for different
organs and tissues [88,89]. Use is made of a computer-aided design (CAD) model [81], where the
patient’s specificity is harmonized with the structural requirements of the scaffold, which results in a
complex, highly accurate, 3D product [88,89].
The process comprises several steps. The first step is to create a CAD model which is transferred to
a file suitable for stereolithography method. In the “pre-processing” step, there is a stereolithography
file digitally sliced into cross-sectional layers. Then one layer is produced, and the printing continues
until the process is completed. The finishing of the structure involves hardening and surface
treatment [88]. Digital data, which facilitate production of precise structures [34] of porous (over
90%) and interconnected scaffolds [89,90], with high reproducibility [88], are received from imaging
techniques such as computer tomography or magnetic resonance. SFFT provides scaffolds with
controlled micro- and macroporous structures [34] in different parts of the same scaffold. Anisotropic
microstructures can be beneficial where multiple cell types are essential. Different SFFT used by
different research groups are evaluated in Table 4 in respect of the advantages and disadvantages
of their technique types. Solid free-form fabrication is a relatively new method. The solution can
be found very quickly, and the scaffold is tailored. Not all SFFT types are used for production of
scaffolds [88–91].
Table 4. Advantages and disadvantages of different SFFT types.
Techniques Materials Advantages Disadvantages
Photo-polymerization of
High accuracy, complex 3D structure
PEG, PEGDA, PPF, PCL, materials, photocurable
Stereolithography (SL) including agents and cells, easy
PDLLA materials, expensive
removal of photopolymer by heating
materials and equipment
High porosity, complete pore
High processing
Thermoplastic polymers interconnectivity, possibility of
Fused deposition temperature, limited
and their composites controlling porosity and size of pores,
modeling (FDM) material range,
(PVA, ABSP400) macro shape control, good compressive
inconsistency in pores,
strength, solvent-free
Complex structure, possibility of High processing
controlling porosity and size of pores temperature, using only
Selective laser sintering Polymer ceramics (PCL,
independently, wide range of powder thermally stable
(SLS) HAp, TCP)
materials, solvent-free, any secondary polymers, limited to
binder system small pore size
Easy process, high porosity, complete
Use of toxic organic
pore interconnectivity, possibility of
Ceramics, polymers, solvent, lack of
3D printing (3D-P) controlling porosity and size of pores
metals mechanical strength,
independently, macro shape control,
limited to small pore size
wide range of materials
PEG: polyethylenglycol, PEGDA: poly(ethylene glycol)diacrylate, PPF: polypropylene fumarate, PVA: polyvinyl
alcohol, ABSP400: acrylonitrile-butadiene-styrene, PCL: polycaprolactone, PDLLA: poly D,L-lactide, HAp:
hydroxyapatite, TCP: tricalcium phosphate.Materials 2019, 12, 568 14 of 25
5.1.5. Microsphere Sintering
In this method, a microsphere composite of ceramic and polymer that is produced by
the emulsion/solvent evaporation technique, is sintered and 3D porous scaffolds are obtained.
The porosity obtained is about 40% and pore diameter is about 90 µm [34]. Sintered microsphere
scaffolds have excellent mechanical properties comparable to cancellous bone. Sintering temperature
and sintering time are the decisive factors. Higher temperature and a longer time give superior fusion
of microspheres, a smaller pore size with lower porosity, and better mechanical properties [92].
Using a solvent (most frequently methylene chloride and acetone) can be another way of
producing sintered microsphere scaffolds [92]. PLGA, as the most often used synthetic polymer,
was used for microsphere sintered scaffold with two types of pores. Larger pores enable blood vessel
and bone tissue ingrowth and smaller surface ones improve the passage of nutrients. Thanks to
smaller surface pores, surface roughness is increased, and cellular attachment and proliferation are
improved [93].
CO2 was used for a subcritical CO2 sintering method. The gas-foaming method uses CO2 currently,
but this method creates a closed-pore structure. It has been claimed that the use of CO2 for producing
scaffolds with microsphere sintering creates interconnective pores with higher porosity. The optimal
range of CO2 pressure was found to be between 15 and 25 bar. The advantage of using CO2 is its
non-toxicity. More research is needed into establishing and setting sintering conditions with CO2 and
adding growth factors without damage occurring [94].
5.1.6. Emulsion Freeze-Drying Method
The emulsion freeze-drying method is based on phase separation including emulsification and
freeze-drying [82], and highly porous scaffolds are produced. The first step is preparation of the
emulsion by homogenization of a polymer in an organic solvent and water. This emulsion is rapidly
cooled down and liquid phases (water and solvent) are removed by freeze-drying. The resultant
pores are close together, but the porosity is higher than 90% and the size of pores is between 20 and
200 µm [81]. The emulsion freeze-drying method could be combined with particulate leaching, sucrose
or sodium chloride can be added in to the emulsion to create porosity. After freeze-drying, particles
can be washed away [95].
5.1.7. Electrospinning Techniques
Electrospinning is a promising versatile technique that uses a high electric field to produce
submicrometer fibers or nanofibers by reduction of surface tension within the polymer fluids.
A solution or melt of a synthetic or natural polymer [96] is injected with an electrical potential
to create a charge imbalance [81], which allows stable, steady deposition of electrospun fibers on any
substrate. Different types of biopolymers (e.g., PCL, PU, collagen) could be electrospun [97]. Due to
the poor stability of natural polymers and the harmfulness of the degradation products of synthetic
polymers, a combination is usually utilized [96].
Electrospinning could generate non-woven matrices with nanoscale features. The thickness of
individual fibers and their orientation could be controlled by type and concentration of the polymer
and by the setting of electrospinning device. The polymeric non-woven nanofiber scaffolds have high
porosity and high surface area [98].
5.1.8. Three-Dimensional Bioprinting
Tissue engineering increasingly exploits 3D printing processes. 3D porous scaffolds need a
consistent and adequate size of well-interconnected pores for cell migration and proliferation. Several
commonly used techniques to produce these 3D scaffolds have been mentioned above, but all have the
same disadvantages: inadequate control of scaffold architecture, pore network and size, and suboptimal
3D scaffolds [99]. These methods are not versatile enough in their processes [100]. 3D-printingYou can also read
